Tissue Engineering Methods

To date, the main reasons for small-diameter vascular graft failure and limitations are due to increased surface thrombogenicity caused by the absence of a functional endothelium and intimal hyperplasia caused by compliance mismatch and chronic inflammation.51 A TEVG can be constructed using various methods, including cell sheets or seeding cells onto biodegradable or decellularized scaffolds.57,62 _{A successful}

vascular graft should be available off-the shelf, durable under long-term implantation, have low inflammatory response, and should be resistant to thrombosis and infection. The vessel wall should also be elastic and match compliance with the native host vessel.51 In order to meet these requirements, the focus on tissue engineering methods are gaining increasing interest. When constructing tissue engineered vascular grafts, there are three major components to be considered: a scaffold with the desired shape and mechanical strength, an adhesive matrix, and cells to be incorporated.63 _{There are a number of design}

criteria to be met; since the construct is meant for supporting blood flow, it must

withstand the pressures exerted by this flow without bursting or permanent deformation. It should also possess appropriate compliance to prevent high stresses around the

anastomosis and have a geometry that will not result in undesirable flow characteristics.59 Blood vessels demonstrate both viscous and elastic behaviors. At low pressures, elastic fibers and contractile SMCs are responsible for vascular mechanical strength and at high pressures, the recruitment of aligned collagen fibers is responsible for the non-linear stress-strain relationship, ultimate tensile strength (UTS), stiffness and elastic modulus. Therefore, native arteries are compliant and elastic at low pressures and strong at high

pressures.64 _{Along with not triggering a negative immune response, the graft should be}

easily implantable and have the ability to be handled, manipulated and sutured.59 Currently, synthetic vascular grafts fabricated from expanded

polytetrafluoroethylene (ePTFE), polyethylene terephthalate (Dacron®) and polyurethane are used, however, due to thrombus formation and compliance mismatch, none of these materials have been suitable for producing grafts less than 6 mm in diameter. For example, only 45% of standard ePTFE grafts are patent as femoropopliteal bypass grafts at 5 years. Designing new vascular replacements is increasingly moving towards

mimicking the native arterial wall to maintain normal physiologic responses of the vascular wall and control thrombosis and inflammation. Collagen and elastin, which are responsible for tensile strength and viscoelasticity of the blood vessel are key structural components. Furthermore, the endothelial lining is another important design

consideration, as the native endothelium is responsible for maintaining a protective, thromboresistant barrier between the blood and tissue as well as controlling vessel tone, platelet activation and leukocyte adhesion.65

The first tissue-engineered blood vessel was developed by Weinberg and Bell in 1986. They cultured bovine endothelial cells, SMCs and fibroblasts in layers of collagen gel supported by a Dacron mesh. Physiological pressures were maintained for only 3-6 weeks, but they could demonstrate the feasibility of a tissue engineered graft with human cells. This achievement pioneered the search for a suitable material for a vascular graft and focused the research on finding appropriate coatings and surface chemical

Since cells in culture cannot organize themselves into a complex three- dimensional structure, the use of a scaffold to provide a template of the construct is a common approach in tissue engineering. This field has seen scaffold construction using a wide range of synthetic and natural matrices with different manufacturing techniques. A variety of synthetic, natural, and hybrid polymer scaffolds have been developed by a number of researchers.59 _{Synthetic materials have been incorporated in vascular graft}

design primarily due to the ease and flexibility of modifying their mechanical properties. Currently, Dacron is the most commonly used for aortic replacement and occasionally as a conduit for femoropopliteal bypass surgery. Dacron grafts are often crimped

longitudinally to increase flexibility, elasticity and kink resistance; however, these properties are usually lost soon after implantation, due to tissue ingrowth. In general, the patency rates of Dacron and ePTFE grafts are similar. Polyurethane is also a viable material for consideration; Nakagawa et al. developed a poly(ether)-urethane graft

reinforced with knitted polyester fibers, which was found to be more durable than ePTFE. Further work has resulted in developing a poly(carbonate-urea)urethane vascular graft that exhibits a compliance similar to human arteries. 65,66

The generally poor patency rates of synthetic polymers have motivated further research in developing ways to functionalize the luminal surface of grafts and target tissue regeneration. Coatings, chemical and protein modifications, and endothelial cell seeding on inert materials have been applied to improve reendothelialization and inhibit inflammation. Cell seeding has proven to have limited success due to difficulties in cell sourcing and attachment, and retention during pulsatile flow conditions. Methods to

promote in situ regeneration of a functional endothelial lining have also been challenging because of chronic inflammatory and prothrombotic responses to the synthetic polymeric materials.65

An alternate method of tissue engineering is the use of biodegradable polymers as scaffolds on which layers of cells are grown. The scaffold degrades and is replaced and remodeled by the extracellular matrix (ECM) secreted by the cells. Polyglycolic acid (PGA)is one such polymer that degrades through the hydrolysis of its ester bonds, and in turn, glycolic acid is metabolized and eliminated as water and carbon dioxide. PGA loses its strength in vivo within 4 weeks and is completely absorbed by 6 months. Although this approach has shown promising results, some drawbacks exist. Cell sourcing is a challenge including long culture periods ranging between 2 and 6 months, and the proliferative capacity of cells isolated from elderly patients is limited. While the mechanical strength of the materials is comparable to that of native vessels, but

compliance mismatch limits long-term patency.65 Nieponice et al. formulated a scaffold by combining a porous biodegradable elastomeric polymer, poly(ester urethane)urea (PEUU), with muscle-derived stem cells using their novel rotational vacuum seeding technique resulting in a cellularized tubular construct. Results showed that the cells were able to proliferate and populate the polymer while in dynamic culture, retained their stem cell features, and produced collagen when stimulated with ascorbic acid. While this study showed promising results, one limitation was the lack of developing a functional endothelial layer to prevent acute thrombosis upon implantation.67

Row et al. incorporated Small Intestinal Submucosa (SIS) and fibrin to engineer vascular grafts. SIS has shown to maintain a porous scaffold structure that enables cell migration and proliferation. In this study, SIS was turned into cylindrical grafts using fibrin glue. These were implanted into the carotid arteries of sheep and exhibited long- term patency and high levels of host cell infiltration. One-month post-implantation, α- SMA expressing cells were dominantly present indicating vascular contractility. They also found that the presence of donor cells was helpful but not necessary for successful host cell infiltration, remodeling or development of vascular function. However, the study determined that the performance of SIS as a small diameter vascular graft is inadequate, likely due to the lack of a functional endothelium.68

Despite all of the advances in tissue engineering incorporating synthetic materials, the limitations for use as small-diameter arterial replacements requires the continued search for optimal methods of engineering materials for this specific application. One approach is the use of decellularized natural matrices; this method takes advantage of the structure and mechanical performance of natural ECM while avoiding any adverse immunological reactions. Decellularized grafts, like synthetic grafts, would be readily available and unlike synthetic grafts, would provide the proper microenvironment for supporting cell infiltration, proliferation and differentiation.50 _{Malone et al. and Lalka et}

al. pioneered the research on tissue decellularization by reporting that implanting cell-free arterial allografts does not result in immunologic alterations. Their research showed that treating the tissues with SDS formed an entirely ECM-based tube with morphologically intact elastin and collagen. The primary take away from these studies was the discovery

that we can focus on reducing allograft/xenograft antigenicity instead of

immunosuppressing the host.69,70 Decellularization is the process of removing antigenic cellular material from the tissue. It can be achieved using a variety of chemical agents such as acids and bases, hypo/hypertonic solutions, detergents and solvents, along with physical methods such as agitation, pressure and abrasion. The native architecture of decellularized tissues along with their diverse structural properties give them many advantages for use as vascular scaffolds. The three-dimensional architecture of the ECM and ECM proteins feature cell-signaling components, which encourage adhesion,

migration, proliferation, and differentiation of host cells. They also have nearly ideal biomechanical properties.51